Magnetic Resonance (MR) Imaging technology is quite commonly used today in larger medical institutions worldwide, and has led to huge benefits in the practice of medicine. A significant factor affecting further use of this versatile imaging technology is the high cost associated with such systems, both for purchase and maintenance.
The high costs associated with design and manufacture of such systems are due in large part to the necessity for a large and very homogeneous static magnetic field, as well as the need for gradient field producing apparatus for imaging with such systems; such large static fields are currently required in order to obtain high image quality and resolution. In such systems the high degree of field homogeneity together with the use of uniform gradient magnetic fields that are applied for imaging purposes, permits a straightforward recovery of the proton or nuclear density distribution in the imaged sample by means of direct Fourier inversion of the received signals. However, the need for substantial homogeneity in the static magnetic field associated with a Magnetic Resonance Imaging system adversely impacts system size and cost.
To achieve a significant reduction in system size and cost of MR Imaging systems, it is useful to be able to image with inhomogeneous static magnetic fields. Typically, current existing commercial methodologies are predicated on the use of a very homogeneous static magnetic field and cannot be used in the presence of inhomogeneous magnetic fields.
Previous attempts at constructing a method for MR Imaging with inhomogeneous static fields have assumed a static magnetic field in homogeneity that is substantially in a single spatial direction, for instance by exciting nuclear spins at or near a measurement surface that corresponds to an isosurface of constant static magnetic field magnitude, with a substantially constant field gradient in a direction normal to the measurement surface, denoted as the z-direction. The latter substantially constant behavior of field gradient implies that the measurement surface is substantially planar. In such schemes, the gradient magnitude is assumed to be quite large in the z-direction, typically of the order of 400 Gauss/cm or 0.04 Tesla/cm. This large gradient magnitude leads to rapid dephasing of the spins after excitation by application of suitable Radio Frequency (RF) pulses. However, this same large gradient can yield short refocusing times when a suitable Radio Frequency pulse sequence is applied in the form of a standard 90-180 spin echo pulse sequence, where the refocusing 180-pulses are applied at a rate proportional to the z-gradient. Repeated spin echoes are thereby produced, and are used to create an averaged or enhanced signal in a time period that is not too large. Suitable intra-slice voxel encoding in the form of applied x- and y-gradients further is used, together with standard MR Imaging and signal inversion techniques to reconstruct an image of the (thin) neighborhood of the measurement surface.
Extensions of such methods include subslicing, wherein an image of a tissue slice is formed by dividing the slice into a number of subslices, the number being determined by the ratio of T1 and T2 relaxation times of the tissue. The subslice images are combined to form an image of the slice. Each subslice is imaged by using a narrow bandwidth Radio Frequency pulse with a frequency equal to the Larmor frequency corresponding to that subslice. Repeated spin echoes are used to create an averaged or enhanced signal in a time period that is not too large. Suitable intra-slice voxel encoding in the form of applied x- and y-gradients further can be used together with standard MR Imaging and signal inversion techniques to reconstruct an image of the subslice. Similar methods are used to image a slice of tissue in the presence of a strong z-gradient by spread spectrum methods to suppress the effect of perturbations in the static field. Repeated spin echoes are used to create an averaged or enhanced signal in a time period that is not too large. Suitable intra-slice voxel encoding in the form of applied x- and y-gradients further can be used together with standard MR Imaging and signal inversion techniques to reconstruct an image of the slice.
The methods mentioned above use a static magnetic field with a pattern of inhomogeneity or gradient in a single spatial direction. Magnets producing such a field pattern in a significantly large volume of interest within a patient for use as a general purpose imaging system would physically be quite large. As a practical matter, while these methods could be used to develop smaller imaging systems to image extremities or small peripheral portions of patient anatomy, they do not easily permit general internal imaging of patient anatomy.
There have been early attempts to develop an imaging probe akin to a catheter with an MR Imaging system utilizing magnets mounted on the probe. The probe is used for “inside-out” imaging to generate an image of a small local region in the anatomy of interest that is external to the probe itself. It produces an image of a wedge-shaped region transverse to the long axis of the probe by using an inhomogeneous static magnetic field produced by the catheter-mounted magnet(s). The strong radial gradient of these magnets, together with catheter-mounted gradient coils that produce a circumferential gradient, are used to phase encode spins in the wedge-shaped region after they are excited by Radio Frequency signals transmitted by transmission coils also mounted on the probe. As in the previously-mentioned methods, repeated spin echoes followed by signal averaging are used to overcome dephasing effects, and standard MR Imaging and signal inversion techniques are then employed to reconstruct an image of the thin wedge-shaped region. By translation of the probe in the longitudinal direction, and rotation of the probe about its axis, an annular cylindrical region surrounding the probe can be imaged. Magnets can also be mounted on an MR imaging probe to create a strong, inhomogeneous local gradient in the presence of a static magnetic field that is produced by a standard MR Imaging system. The local imaging techniques described earlier in this paragraph can generate a high resolution local image of a local region surrounding the probe, which can then be superimposed on a larger field-of-view image generated by the standard MR Imaging system.
The imaging methods discussed in the previous paragraph are specifically designed for use in external imaging in a local region around a probe. They rely on the use of repeated spin echoes and averaging in the presence of a strong locally produced gradient. Furthermore, the use of an imaging probe is invasive. As such, these techniques do not permit general, non-invasive imaging within a wide field-of-view.